Method and apparatus for determining respiratory system resistance during assisted ventilation

ABSTRACT

Method and apparatus are described for determining respiratory system resistance (R) in a patient receiving gas from a ventilator. A negative pulse in the pressure and/or flow output of the ventilator during selected inflation cycles is generated and Paw, V dot and V are measured at a point (T 0 ) near the beginning of the pulse, at a point (T 1 ) near the trough of the negative pulse and at a point (T −1 ) preceding T 0 . The value of R is calculated from the difference between Paw. V dot and V at T O  and at T I  and where the change in patient generated pressure (Pmus) in the interval T O −T I  is estimated by extrapolation from the different between Paw, V dot and V and T 0  and at T −I , in accordance with Equation 8.

FIELD OF INVENTION

[0001] This invention relates to mechanical ventilation, and inparticular, to assisted ventilation and the determination of respiratorysystem resistance.

BACKGROUND TO THE INVENTION

[0002] There are currently no reliable, clinically available,non-invasive means to estimate respiratory resistance (R) duringinspiration in mechanically ventilated patients who have spontaneousrespiratory efforts. Calculation of resistance requires knowledge of theforce applied to the respiratory system which, in such patients,includes a component related to pressure generated by respiratorymuscles (Pmus). This component continuously changes during the inflationphase and cannot be estimated without prior knowledge of respiratorymechanics. Furthermore, to isolate the component of total appliedpressure that is dissipated against resistance (P_(res)), it isnecessary to subtract the pressure used against the elastic recoil ofthe respiratory system. This requires knowledge of passive respiratoryelastance (E) which is also difficult to determine in the presence ofunquantifiable Pmus. At present, therefore, R can be reliably estimatedonly by use of esophageal catheters, which add another invasiveintervention to already much instrumented patients, or by elimination ofrespiratory muscle pressure output with paralysis, or hyperventilation(controlled mechanical ventilation, CMV). The latter entails additionalpersonnel time and does not lend itself to frequent determination of R.To the extent that R is a highly dynamic property that may changefrequently, due to secretions or changes in bronchomotor tone,availability of continuous estimates of R may be helpful in the clinicalmanagement of such patients. Thus, changes in R can be rapidlyidentified and dealt with. Furthermore, this information makes itpossible to adjust the level of assist according to the prevailing Rvalues, a feature that is of particular utility in pressure assistedmodalities of ventilatory support (Pressure Support Ventilation,Proportional Assist Ventilation).

[0003] In U.S. Pat. No. 5,884,622 (Younes), assigned to the assigneehereof, an approach is described to determine resistance under similarconditions, namely in assisted ventilation. This prior approach consistsof applying at least two different types of transient changes in flow inthe course of the inflation phase of the ventilator. The changes inairway pressure (Paw), flow ({dot over (V)}), and volume (V) duringthese transient flow changes are compared with the time course of thesevariables in unperturbed breaths. While this approach is capable ofproviding accurate information about R, it has several limitations.First, because of considerable breath-by-breath variability in the timecourse of Paw, {dot over (V)} and V in spontaneous unperturbed breaths,it is necessary to average large numbers of perturbed and unperturbedbreaths in order to arrive at the real change that occurred during theperturbation. Accordingly, information about resistance is delayed untila sufficiently large number of observations has been averaged.Furthermore, for the same reasons, any true change in patient'sresistance is not detected in a timely way. Second, this approachrequires at least two different kinds of perturbations. Because, asindicated earlier, a large number of observations is required with eachperturbation, this requirement delays the acquisition of reliableinformation further. Third, the need to average large numbers of breathsand a large number of data points from each breath, greatly increasesthe computing and storage requirements of the computer used to processthe information to provide the value of R. This requirement adds furtherstrain on the extensive and highly complex operations carried out bymodem, computer controlled ventilators.

SUMMARY OF INVENTION

[0004] The method and apparatus described in detail herein in accordancewith the present invention, represent a considerable simplification ofthe approach proposed by Younes in U.S. Pat. No. 5,884,622. As indicatedabove, the main obstacle to determining respiratory resistance duringassisted ventilation is the uncertainty about what happens to Pmusduring interventions in the course of the inflation phase of theventilator. The comparison between perturbed and unperturbed breaths wasthe approach used in U.S. Pat. No. 5,884,622. By contrast, in accordancewith the present invention, the behavior of Pmus during the interventionis predicted from estimates of the change in Pmus in the intervalimmediately preceding the intervention. In this manner, all the requiredinformation necessary to determine R can be obtained from a singleintervention in a single breath. This approach greatly reduces thecomputational requirements necessary to determine R, and the timerequired to obtain information that is clinically useful, such as inassisted ventilation

[0005] In accordance with the present invention, respiratory resistance(R) is determined while allowing for the presence of pressure generatedby respiratory muscles (Pmus) but without requiring knowledge of itsactual value or an accurate value of passive respiratory elastance (E).

[0006] In accordance with one aspect of the present invention, there isprovided a method of determining respiratory system resistance (R) in apatient receiving gas from a ventilatory assist device (ventilator),comprising estimating the flow rate ({dot over (V)}) and volume (V) ofgas received by the patient from the ventilator, estimating pressurenear the airway of the patient (Paw), generating a signal that resultsin a step decrease (negative pulse) in the pressure and/or flow outputof the ventilator during selected inflation cycles, measuring Paw, {dotover (V)} and V at a point (T_(O)) near the beginning of the pulse(Paw_(O), {dot over (V)}, V_(O)), at a point (T_(I)) near the trough ofthe negative pulse (Paw₁, {dot over (V)}_(I), V₁), and at a point (T⁻¹)preceding T_(O) but after the onset of inspiratory effort (Paw⁻¹, {dotover (V)}⁻¹, V⁻¹); and calculating the value of resistance (R) from thedifferences between Paw, {dot over (V)} and V at T_(O) and at T_(I) andwhere the change in patient generated pressure (Pmus) in the intervalT₀→T₁ (ΔPmus ( T_(O)→T_(I))) is estimated by extrapolation, from thedifferences between Paw, {dot over (V)} and V at T_(O) and at T⁻¹, inaccordance with equation (8).

[0007] As described in more detail below, the present invention includesmodifications to the method as alternative steps to determining R.

[0008] In accordance with another aspect of the present invention, thereis provided an apparatus which interfaces with ventilatory assistdevices (ventilators) determining respiratory system resistance (R),comprising a flowmeter, with associated electronic circuitry, thatestimates the flow rate ({dot over (V)}) and volume (V) of gas receivedby a patient, a pressure sensor that estimates pressure near the airwayof the patient (Paw), and electronic circuitry which receives the Paw,{dot over (V)} and V signals from above mentioned circuitry and which isalso connected to the control system of the ventilator, comprising:

[0009] circuitry that generates an output that results in a stepdecrease (negative pulse) in the pressure and/or flow output of theventilator during selected inflation cycles;

[0010] circuitry that measures Paw, {dot over (V)} and V at a point(T_(O)) near the beginning of the pulse (Paw_(O), {dot over (V)},V_(O)), at a point (T_(I)) near the trough of the negative pulse (Paw₁,{dot over (V)}₁, V₁), and at a point (T⁻¹) preceding T_(O) but after theonset of inspiratory effort (Paw⁻¹, {dot over (V)}⁻¹, V⁻¹);

[0011] circuitry to calculate the value of resistance (R) from thedifferences between Paw, {dot over (V)} and V at T_(O) and at T₁ andwhere the change in patient generated pressure (Pmus) in the intervalT_(O)→T_(I) (ΔPmus (T_(O)→T_(I))) is estimated, by extrapolation, fromthe differences between Paw, {dot over (V)} and V at T_(O) and at T⁻¹,in accordance with equation (8).

[0012] As described in more detail below, the present invention includesmodifications to the apparatus as alternative combinations of elementsto determine R.

BRIEF DESCRIPTION OF DRAWINGS

[0013]FIG. 1 shows a tracing of airway presence (Paw), flow and volumeshowing a negative pulse and the three times at which measurements aretaken;

[0014]FIG. 2 is a schematic representation of apparatus for carrying outthe method in accordance with a preferred embodiment of the invention;and

[0015]FIG. 3 shows schematically the various elements of a microcontroller used in connection with the apparatus of FIG. 2.

GENERAL DESCRIPTION OF THE INVENTION

[0016] According to the equation of motion, the total pressure appliedto the respiratory system (P_(appl)) is dissipated against elastic,resistive and inertial opposing forces. Thus:

P _(appl) =P _(el) +P _(res) +P _(iner) where:

[0017] P_(el) is elastic recoil pressure and is given by the product ofvolume above passive functional residual capacity (FRC) (V) andelastance (E); P_(el)=V.E,

[0018] P_(res) is the pressure dissipated against resistance and isgiven by the product of flow ({dot over (V)}) and R; P_(res)={dot over(V)}.R, and,

[0019] P_(iner) is the pressure dissipated against inertia and is givenby the product of flow accelaration (the rate of change in flow inl/sec²; {umlaut over (V)}) and inertia (I). Because I of the respiratorysystem is very small (≈0.02 cmH₂O/l/sec²), P_(iner) can be ignored solong as measurements are made at relatively low {umlaut over (V)} (e.g.<10 l/sec²). In mechanically ventilated patients, {umlaut over (V)} mayexceed this level only in the first about 100 to 200 msec of theinflation phase during volume cycled and high level pressure supportventilation (PSV). Accordingly, by avoiding measurements in this region,the equation solved to calculate P can be simplified by neglectingP_(iner).

[0020] During assisted ventilation, P_(appl) is made up of twocomponents, one provided by the ventilator (Paw) and one provided by thepatient (Pmus). Thus, P_(appl)=Paw+Pmus. With this equation and earlierconsiderations, the equation of motion can be rewritten and rearrangedas follows as equation (1):

{dot over (V)}.R=Paw+Pmus−V.E  (1)

[0021] To the extent that Pmus at a given instant is not known, accurateelastance values may not be available and V, relative to passive (FRC),is also not known (in view of possible dynamic hyperinflation or activereduction in volume below FRC by expiratory muscles), it is not possibleto solve for R using a set of measurements made at one point during theinflation phase. For this reason, any approach to measure resistanceduring inflation in such patients must involve measurements at more thanone point, having different flow values, as described herein.

[0022] In one aspect of the present invention, Paw (and hence flow) israpidly reduced (negative pulse) during the inflation phase (FIG. 1).Primary measurements of Paw, {dot over (V)} and V are made at or nearthe point where Paw and flow begin declining (T_(O)) and at or near thetrough of pressure during the negative pulse (T_(I)). These two samplingpoints are chosen because ΔP/Δt and ΔV/Δt are minimal. In this fashion,inertial forces can continue to be ignored. More importantly, errorsrelated to differences in delay and frequency response of the pressureand flow measuring systems can be avoided. This advantage isparticularly relevant since, in modern ventilators, Paw and patient floware not measured directly near the ET tube but are estimated from remotesites and, hence, the signals may be subject to different delays andresponse characteristics. Even minor differences in these properties cancause serious errors when Paw and patient flow are changing rapidly (forexample, during the declining phase of the pulse).

[0023] Equation 1 can be written for T_(O) and T_(I) as follows, asequations (2) and (3):

{dot over (V)} _((O)) .R=Paw _((O)) +Pmus _((O)) −V _((O)) .E  (2)

{dot over (V)} _((I)) .R=Paw _((I)) +Pmus _((I)) −V _((I)) .E  (3)

[0024] Subtracting equation 3 from equation 2 yields equation (4):

R({dot over (V)} _((O)) −{dot over (V)} ₍₁₎)=(Paw _((O)) −Paw_((I)))+(Pmus _((O)) −Pmus _((I)) )−E(V _((O)) −V _((I)))  (4)

[0025] Rearranging equation (4) to solve for R:

R=[(Paw _((O)) −Paw ₍₁₎)+ΔPmus(T _(O) →T ₁)−E(V _((O)) −V ₍₁₎)]/({dotover (V)} _((O)) −{dot over (V)} ₍₁₎)

[0026] In theory, if the time interval between T_(O) and T_(I) (i.e. Δt)is infinitely small, the differences in Pmus and in volume can beignored and ΔPaw becomes ΔPres. In practice, however, during mechanicalventilation it is not possible to instantly reduce flow from one valueto another relatively stable value (i.e. at which ΔP/Δt and Δ{dot over(V)}/Δt are acceptably small). Even if flow exiting the ventilator isaltered suddenly, a finite time must elapse before the flow to thepatient stabilizes at the new value in view of continued flow from thetubing to patient in the process of decompression of the circuit. Δt,therefore, cannot be made short enough to ignore changes in Pmus betweenT_(O) and T_(I) and ΔPmus in this interval has to be accounted for.

[0027] In this aspect of the present invention, the change in Pmusbetween T_(O) and T_(I) is estimated by assuming that Pmus changes inthis time range at the same rate as in the period immediately precedingT_(O). This is not an unreasonable assumption if the time intervalbetween T_(O) and T₁ is relatively brief (for example, approximately 100msec). The rate of change in Pmus immediately before the pulse isestimated by sampling Paw, {dot over (V)} and V at a point shortlybefore T₀(for example, 100 msec prior to T_(O)) (T⁻¹), (FIG. 1). Thefollowing two equations (5) and (6) provide estimates of Pmus at T_(O)and T⁻¹ respectively and represent rearrangement of the equation ofmotion (equation (1)):

Pmus _(O) =V _(O) E+{dot over (V)} _(O) .R−Paw ₀  (5)

Pmus ⁻¹ =V ⁻¹ .E+{dot over (V)} ⁻¹ .R−Paw ⁻¹  (6)

[0028] Subtracting equation 6 from equation 5 and dividing by Δt−1 (timebetween T_(O) and T_(−I)) gives ΔPmus/Δt in the interval (Δt⁻¹) prior toT_(O) according to equation (7):

ΔPmus/Δt=(1/Δt _(−I))[E(V _(O) −V ⁻¹)+R({dot over (V)}_(O) −{dot over(V)} ⁻¹)−Paw _(O) +Paw ⁻¹]  (7).

[0029] Assuming that Pmus changes at the same rate between T_(O) andT_(I), the change in Pmus between these two points is given by:

(Pmus _(O) −Pmus _(I))=−Δt _(I)  [equation 7]

[0030] where Δt_(I) is the time interval between T_(I) and T₀.

[0031] Substituting [−Δt (equation 7)] for (Pmus_(O)−Pmus₁) in equation4 and rearranging provides equation (8):

R=[(Paw _(O) −Paw ₁+(Δt/Δt ⁻¹)(Paw ₀ −Paw ⁻¹))−E(V _(O) −V ₁+(Δt/Δt_(−I)) (V _(O) −V _(−I)))]/({dot over (V)} _(O) −{dot over (V)}_(−I)))  (8)

[0032] The only unknown in the numerator of equation 8, which is theestimate of Pres, is E. However, unlike the case in equation 4, thedifference in V between T_(O) and T₁ is now reduced by the term(Δt_(I)/Δt⁻¹) (V_(O)−V_(I)). If the pulse is initiated during the risingphase of flow (e.g. FIG. 1), average {dot over (V)} in the intervals T⁻¹to T_(O) and T_(O) to T₁ will not be substantially different and, giventhat the time intervals between T_(O) and T₁ and between T_(O) and T₁are quite small (ca 0.1 sec), the entire volume term is reduced tonearly zero. Under these conditions, any errors in estimating E shouldresult in very minor errors in estimating Pres, and hence resistance(R), and, in the absence of a known value of E, a default value,representing, for example, average E in ventilator dependent patients,can be used without much risk of significant errors. It should also benoted that, because all volume points are obtained from the same breath,differences between any two volume values represent differences inabsolute volume, relative to passive FRC. As a result, offsets ofvolume, relative to passive FRC, at the beginning of the breath becomeirrelevant.

[0033] The above derivation of equation 8 entails the assumption thatthe value of R is constant or, specifically, that R is independent offlow and volume. In reality, R may vary with flow, particularly inintubated patients, if only because the resistance of the endotrachealtube increases with flow. Likewise, R may be dependent on lung volume insome patients. Equation 8 can be adapted to allow for R being flowand/or volume dependent. The number of mathematical functions that canbe used to characterize flow or volume dependence of R is infinite. Itwould be impractical to provide formulations of equation 8 that allowfor all conceivable mathematical descriptions of flow and/or volumedependence. Rather, one example will be illustrated which represents themost widely accepted behavior of R in mechanically ventilated intubatedpatients, namely that R is minimally (or not at all) affected by volumebut that it increases with flow according to Rohrer's equation (R=K₁+K₂{dot over (V)}). It is recognized, however, that any individual withmodest mathematical skills can utilize the same information obtained inthis aspect of the present invention (i.e. Paw, V and {dot over (V)},measured at T_(O), T₁ and at points preceding T_(O)) to derive thepressure-flow relation where mathematical functions other than Rohrer'sequation are assumed to apply.

[0034] The following equations (2a to 8a) correspond to equations 2 to 8above after making appropriate modifications to allow for R to increasewith flow according to Rohrer's equation (R=K₁+K₂ {dot over (V)}):

K ₁ {dot over (V)} _(O) +K ₂ {dot over (V)} _(O) ² =Paw _(O) +Pmus _(O)−V _(O) .E  (2a)

K ₁ {dot over (V)} ₁ +K ₂ {dot over (V)} _(O) ² =Paw ₁ +Pmus ₁ −V ₁.E  (3a)

K ₁({dot over (V)} _(O) −{dot over (V)} ₁)+K ₂({dot over (V)} _(O) ²−{dot over (V)} _(I) ²)=(Paw _(O) −Paw _(I))+(Pmus _(O) −Pmus _(I))−E(V_(O) −V _(I))  (4a)

Pmus _(O) =V _(O) .E+{dot over (V)} _(O) .K ₁ +V _(O) ² .K ₂ −Paw_(O)  (5a)

Pmus _(I) =V ⁻¹ .KE+{dot over (V)} ⁻¹ .K ₁ +V ⁻¹ ² .K ₂ −Paw ⁻¹  (6a)

ΔPmus/Δt=(1/Δt _(−I))[E(V _(O) −V ₁)+K ₁({dot over (V)} _(O) −{dot over(V)} ⁻¹)+K ₂({dot over (V)} _(O) ² −{dot over (V)} ⁻¹ ²)−Paw _(O) +Paw⁻¹]  (7a)

K ₁({dot over (V)} _(O) −{dot over (V)} ₁+(Δt _(I) /Δt _(−I))({dot over(V)} _(O) −{dot over (V)} ⁻¹))+K ₂({dot over (V)} _(O) ² −{dot over (V)}₁ ²+(Δt1/Δt−1)({dot over (V)} _(O) ² −{dot over (V)} ⁻¹ ²)=(Paw _(O)−Paw ₁+(Δt/Δt _(−I))(Paw _(O) −Paw ⁻¹))−E(V _(O) −V ₁+(Δt/Δt _(−I))(V_(O) −V ⁻¹))  (8a)

[0035] From each applied pulse, an equation of the form of equation (9)accordingly results:

K ₁ .X+K ₂ .Y=Z  (9)

[0036] where X is the flow term (first bracketed term to left ofequation 2a), Y is the {dot over (V)}² term (second bracketed term inequation 8) and Z is the Pres term (right side of equation 8). To obtainZ, a known value of E is used or, in the absence of this information, adefault value (e.g. 28 cmH₂O/l, representing average E in mechanicallyventilated patients (personal observations), may be used. Resistance canbe obtained from the above equation (9) in one of several ways. Some ofthese are listed below:

[0037] 1) K₂ is initially assumed to be zero and resistance is estimatedfrom Z/X. The resistance value obtained in this fashion represents theslope of the P {dot over (V)} relation between {dot over (V)}_(O) andweighted average of {dot over (V)}₁ and {dot over (V)}⁻¹. If {dot over(V)}₁ and {dot over (V)}⁻¹ are not substantially different (e.g. FIG.1), R calculated in this fashion can be assumed to represent the slopeof the P {dot over (V)} relation between {dot over (V)}_(O) and either{dot over (V)}_(I) or {dot over (V)}⁻¹ or the mathematical average ofthe two. It can be shown, using Rohrer's equation, that the slope of theP {dot over (V)} relation between any two flow points (incrementalresistance, IR) is the same as the resistance at a flow corresponding tothe sum (flow-sum) of the two flow points (in this case ({dot over(V)}_(O)+{dot over (V)}₁)). With this treatment, R is reported asresistance at a specific flow (i.e. flow sum).

[0038] 2) If a range of flow-sum values is obtained in successivepulses, either spontaneously or by design, a range of IR values willalso result. A regression between IR (dependent variable) and flow-sumwill result in a significant correlation if a sufficiently wide range offlow-sum is present. The intercept of this regression is K_(I) and theslope is K₂. These can be reported as such. Alternatively, resistancemay be reported as the sum of K_(I) and K₂, which is resistance at astandard flow of 1.0 l/sec. This has the advantage that changes inreported resistance reflect real changes in resistance whereas withapproach #1, alone, the reported resistance may change simply becauseflow is different.

[0039] 3) The values of K_(I) and K₂ can be derived from the results oftwo pulses having different X and Y values, or by regression analysis ofthe results of multiple pulses displaying a range of X and Y values. Theprocedure of applying pulses can be deliberately planned to result in awide range of X and Y values in order to facilitate this analysis. Forexample, pulses may be initiated at different flow rates, so that {dotover (V)}_(O) is variable, and/or the decrease in {dot over (V)} duringthe pulse can also be deliberately varied, to result in a range of {dotover (V)}₁.

[0040] 4) In the absence of reliable, directly determined K_(I) and K₂values, following approach #2 above, K₂ can be assumed to equal K₂ ofthe endotracheal tube (ET) and equation 9 is solved for K₁. Thus,K₁=(Z−(Y.K₂ET))/X. The K₂ values of clean ET tubes of different sizesare widely available. Resistance can be reported as K₁+K₂ ET, reflectingresistance at a standard flow of 1.0 l/sec. The resistance so reportedmay differ from actual resistance at 1 l/sec to the extent that actualK₂ET may differ from the assumed K₂ of a clean tube AND the flow atwhich R estimates are made are different from 1.0 l/sec. The error inestimated resistance (at 1 l/sec), if actual K₂ (K₂ actual) is differentfrom assumed K₂ is given by R_(error)=(K₂ actual−K₂ assumed) (1−Y/X). Itcan be seen that the error in estimating R at 1 l/sec using an assumedK₂ is a fraction of the difference between the actual and assumed K₂value.

[0041] Potential Sources of Errors and Approaches to Minimize SuchErrors:

[0042] 1) Measurement noise: In mechanically ventilated patients, thePaw and {dot over (V)} signals are subject to noise from multiplesources. These include airway secretions, cardiac artifacts, liquid inthe tubing and oscillations or vibrations in the flow delivery system ofthe ventilator. The noise in the Paw signal may be in phase or out ofphase with that in the {dot over (V)} signal depending on the source ofnoise and the frequency response of the two measuring systems. Out ofphase noise has a greater impact on estimated R particularly if thecritical measurements (e.g. at T₀, T_(I) and T⁻¹) are obtained fromdiscrete points of unfiltered signals. Such noise results in anincreased random variability of estimated R in successive measurements.A more systematic error may result if the pulse is programmed to beginwhen a certain flow is reached. Here, there is an increased probabilitythat the pulse will begin on the upswing of a positive flow artifact.

[0043] Errors related to measurement noise can be reduced by a varietyof approaches:

[0044] a) The most effective approach is to insure that the change inflow produced by the intervention (i.e. change in flow between T_(o) andT₁) is large relative to the amplitude of the noise.

[0045] b) Elimination of sources of noise to the extent possible.

[0046] c) Critical filtering of the Paw and {dot over (V)} signals.

[0047] d) To minimize systematic errors, the pulse should preferably notbegin when a fixed level of flow is reached (see above).

[0048] e) Averaging the resistance results obtained from a number ofpulses.

[0049] 2) Difference in response characteristics of Paw and {dot over(V)} measuring systems:

[0050] Difference in response characteristics of the measuring systemscauses the peak and trough of the measured pressure to occur atdifferent times relative to the flow signal even if the peaks andtroughs of the two signals were, in reality, simultaneous. If T_(O) istaken as the time of peak Paw, flow at T_(O) will underestimate realflow, and vice versa. Also, such differences convert the relativelyinnocuous in-phase oscillations originating from ventilator flowdelivery systems to potentially more serious out-of-phase oscillationsin Paw and flow. To minimize the impact of these differences, the phaselag between the Paw and flow measuring systems should be as short aspossible over the frequencies of interest. In addition, the pulse can bedesigned to avoid sharp peaks and troughs.

[0051] 3) Errors related to extrapolation of the Pmus trajectory:

[0052] These are potentially the most serious particularly whenrespiratory drive, and hence ΔPmus/Δt, is high. The proposed approachinvolves the assumption that ΔPmus/Δt during the pulse is the same asΔPmus/Δt over a finite period prior to the pulse. This assumption can bein error for a variety of reasons. These, and possible ways to minimizethese potential errors, are discussed below:

[0053] a) Termination of inspiratory effort (neural T_(i)) during thepulse: This can potentially produce the largest errors in estimated R.Thus, assume that ΔPmus/Δt prior to T_(O) is 40 cmH₂O/sec and Δt_(I)(i.e. T_(I)−T_(O)) is 0.15 sec. The estimated increase in Pmus betweenT_(O) and T₁ would be 6 cmH₂O. If, however, neural T_(i) ends nearT_(O), Pmus will decrease instead of increasing. Because the rate ofdecline in Pmus during neural expiration is fastest soon after the endof neural T_(i), the decrease in Pmus may actually be greater than theassumed extrapolated increase, with the error in estimated ΔPmusbeing >12 cmH₂O. It can be seen from equation 4 that this conditiontranslates into an error of corresponding magnitude in estimated Pres.If the difference between {dot over (V)}_(O) and {dot over (V)}₁ is 0.4l/sec, this error would translate into an error of >30 cmH₂O/l/sec inestimated resistance.

[0054] Because of the potentially large magnitude of this error, it isnecessary to insure that peak Pmus (end of T_(i)) does not occur betweenT_(O) and T₁. This condition is easy to accomplish during ProportionalAssist Ventilation (PAV). In this mode, the end of ventilator cycle isautomatically synchronized with patient effort and is constrained tooccur during the declining phase of Pmus . So long as pulses are notdelivered in the last fraction (ca 30%) of ventilator T_(I), one isassured that T_(i) termination did not occur within the pulse. Withpressure support ventilation (PSV) and assisted volume-cycledventilation, such synchrony is not assured, however, and T_(i) mayterminate at any point within or even beyond the inflation phase. T_(i)termination may occur, per chance, during some of the pulses resultingin errors of differing magnitudes depending on ΔPmus/Δt prior to thepulse, the point within the pulse at which T_(i) terminated, the rate ofdecline in Pmus beyond the peak, and the difference in flow betweenT_(O) and T_(I). Considerable variability may occur between the resultsof different pulses. For this reason, application of this approachduring PSV and volume cycled ventilation may produce less reliableresistance values.

[0055] b) Shape of the rising phase of Pmus: The rate of rise of Pmusduring the rising phase is not constant. Differences between ΔPmus/Δt inthe interval T_(O) to T_(I)(i.e. Δt_(I)) and T_(−I) to T_(O)(i.e.Δt_(−I)) causes errors in estimated R for the same reasons discussedunder (a) above. A Δt of approximately 0.1 sec is both feasible andconsistent with minimal errors related to response characteristics ofthe measuring systems. It is unlikely that an important change inΔPmus/Δt would occur over this brief time interval, provided allmeasurement points (i.e. T_(O), T_(I), T⁻¹) occur during either therising or declining phase of Pmus. What needs to be avoided is theoccurrence of T_(i) termination between T⁻¹ and T₁ and this can beaccomplished by insuring that pulse application occurs either early inthe inflation phase or very late in the inflation phase in the PAV mode.In this mode, there is assurance that pulses applied in the first 50% ofthe inflation phase occur, in totality (i.e. T⁻¹ to T₁), on the risingphase of Pmus while pulses applied very near the end of the inflationphase will occur in totality on the declining phase of Pmus. In eithercase, there is little likelihood of a major change in ΔPmus/Δt over thebrief period of the pulse and extrapolation from one segment to thenext, within the brief pulse period, should not result in significanterrors.

[0056] c) Behavioral responses: The change in Pmus following theinitiation of the pulse may deviate dramatically from that expected fromthe preceding time interval if the patient perceives the pulse andreacts behaviorally to it. The minimum latency for behavioral responsesto changes in Paw and flow is approximately 0.2 sec, even in very alertnormal subjects. It follows that errors related to perception of thepulse, with consequent behavioral responses, can be avoided ifmeasurements are restricted to the approximately 0.2 sec interval afterinitiation of the pulse. Behavioral responses, however, can occurwithout perception if the change is anticipated. For example, if aperturbation occurs regularly every 5 breaths, the patient may alterhis/her respiratory output every fifth breath, even before the pulse isinitiated. The occurrence of anticipatory responses can be minimized byrandomizing the order of pulse applications.

[0057] d) Non-behavioral neuromuscular responses to changes in flow: Therapid reduction in flow in the course of an ongoing inspiratory effortmay, theoretically, elicit reflex changes in neural output with muchshorter latencies than behavioral responses. In addition, the change inflow and, consequently, in time course of volume, may elicit changes inPmus, independent of changes in electrical activation, through theoperation of the intrinsic properties of respiratory muscles(force-length and force-velocity relations). An important contributionfrom either of these responses following the onset of the pulse (betweenT_(O) and T_(I)) could alter the time course of Pmus relative to thecourse predicted from the pre-pulse interval and introduce errors inestimated Pres. Based on experimental results, these effects are likelyto be small if the change in flow is modest (for example, <1 l/sec) andthe intervention is carried out early in the inflation phase where Pmusis relatively low.

[0058] e) Pmus noise: Noise in the Pmus signal can introduce errors whenthe change in Pmus over a relatively brief period (for example, 0.1 sec)is used to estimate the change in a subsequent interval. Pmus noise canbe real or artifactual. Tracings of P_(di) (transdiaphragmaticpressure), for example, often have a jagged rising phase. Furthermore,when Pmus is estimated from P, {dot over (V)} and V, as opposed to beingmeasured, independent noise in the pressure and flow signals (forexample, cardiac artifact, secretions . . . etc) can introduce noise inestimated Pmus, even if the true rising phase of Pmus is smooth. Theimpact of Pmus noise on estimated resistance is the same whether thenoise is real or artifactual. Random noise in the Pmus signal may beexpected to increase variability in measured resistance values, reducingthe reliability of information obtained from single pulses. Thiscondition can be dealt with by averaging the results of severalobservations over a number of breaths.

[0059] Furthermore, the impact of Pmus noise can also be reduced byusing a relatively large change in flow between T_(O) and T_(I).

[0060] Alternative Approaches to Calculation of Resistance Using thePulse Technique:

[0061] 1) Estimation of the Change in Pmus During the Pulse Using anInterpolation Approach:

[0062] In the above description, the change in Pmus between T_(O) and T₁(i.e. ΔPmus (T_(O)→T₁)) was estimated by extrapolation of the Pmustrajectory in the interval T⁻¹ to T_(O). An alternative approach is toestimate ΔPmus (T_(O)→T₁) by interpolation between two points, onebefore (for example, at T_(O)) and one after the trough of Paw (T₂). Inthis case, Paw, {dot over (V)} and V are measured at T_(O) (i.e.Paw_(o), {dot over (V)} and V_(O)) and at T₂ (i.e. Paw₂, {dot over (V)}₂and V₂) in addition to at T₁. T₂ should preferably be chosen at a point,after T₁, where ΔPaw/Δt and/or Δ{dot over (V)}/Δt are very small tominimize inertial forces. With this alternative approach, equation 8 canbe written as follows:

R=[(Paw _(O) −Paw ₁+(Δt _(I) /Δt ₂)(Paw ₂ −Paw _(O)))−

E(V _(O) −V ₁+(Δt _(I) /Δt ₂)(V ₂ −V _(O)))]/

({dot over (V)} _(O) −{dot over (V)} _(I)+(Δt _(I) /Δt ₂)({dot over (V)}₂ −{dot over (V)} _(O)))  (8 inter)

[0063] where Δt₂ is the interval between T₂ and T_(O). Equation 8a canbe written as follows:

K ₁({dot over (V)} _(O) −{dot over (V)} ₁+((Δt ₁ /Δt ₂)({dot over (V)} ₂−{dot over (V)} _(O))))+K ₂({dot over (V)} _(O) ² −{dot over (V)} ₁²+((Δt ₁ /Δt ₂)({dot over (V)} ₂ ² −{dot over (V)} ₀ ²)))

=(Paw _(O) −Paw ₁+((Δt _(I) /Δt ₂)(Paw ₂ −Paw _(O))))−

E(V _(O) −V ₁+((Δt _(I) /Δt ₂)(V ₂ −V _(O))))  (8a inter)

[0064] There are advantages and disadvantages to the interpolationapproach, relative to the extrapolation approach described earlier. Themain advantage is that, in principle, estimating an unknown value byinterpolation between values before and after the unknown value is moreaccurate than estimating the unknown value through extrapolation of datapoints which are all occurring before or after the unknown value. Thepractical disadvantages in this particular application, however, arethat point T₂ occurs beyond the pulse intervention and, as well, laterin inspiration. Pmus at T₂ may thus be corrupted through behavioral orreflex responses to the preceding intervention, and by the greaterlikelihood that termination of inspiratory effort with precipitousdecrease in Pmus, may occur prior to T₂,

[0065] 2) Combined use of the extrapolation and interpolationtechniques:

[0066] R can be estimated using both the extrapolation technique(equation 8 or 8a) and the interpolation technique (equation 8 inter and8a inter) and the results of the two approaches may be averaged using asuitable averaging technique.

[0067] While either the interpolation approach or the combined approachmay be used in preference to the extrapolation technique, my practicalexperience favors the extrapolation technique. Thus, it was found instudies on 67 ventilator dependent patients that the results of theextrapolation approach are in closer agreement with results obtainedduring controlled ventilation than the results of the other twoalternative approaches.

[0068] 3) Use of back extrapolation, instead of forward extrapolation,of Pmus:

[0069] The change in Pmus between T_(O) and T₁ can be estimated by backextrapolation of data from a period following T₁. Thus, Paw, {dot over(V)} and V are measured at T₂ (see alternative approach #1 above).Equation 8 and 8a are modified to reflect these sampling points asfollows:

R=[(Paw _(O) −Paw ₁+(Δt _(I) /Δt ₂)(Paw ₂ −Paw _(I)))−

E(V _(O) −V ₁+(Δt _(I) /Δt ₂)(V ₂ −V ₁))]/

({dot over (V)} _(O) −{dot over (V)} ₁+(Δt ₁ /Δt ₂)({dot over (V)} ₂−{dot over (V)} ₁))  (equation 8 (bextra)

[0070] and

K ₁({dot over (V)} _(O) −{dot over (V)} ₁+(Δt _(I) /Δt ₂)({dot over (V)}₂ −{dot over (V)} ₁)))+K ₂({dot over (V)} _(O) ² −{dot over (V)} ₁²+((Δt ₁ Δt ₂)({dot over (V)} ₂ ² −{dot over (V)} ₁ ²)))

=(Paw _(O) −Paw ₁+((66 t ₁ /Δt ₂)(Paw ₂ −Paw _(I))))−

E(V _(O) −V ₁+((Δt _(I) /Δt ₂)(V ₂ −V _(I))))  (equation 8a (bextra)

[0071] 4) ΔPmus/Δt prior to T_(O) can be estimated from values obtainedat two points both of which occurring before T_(o) (e.g. T⁻¹ and T⁻²).Although feasible and should provide reasonably accurate results, it haslittle advantage over the use of T_(O) and only one preceding point,while adding more computation complexities.

[0072] 5) Use of regression analysis to estimate ΔPmus/Δt prior T_(O) orbeyond T_(I): The extrapolation approach described above utilizesmeasurements at only two points (e.g. T_(O) and T⁻¹) to estimateΔPmus/Δt(T_(O)→T₁). Although computationally much more intensive,ΔPmus/Δt prior to the onset of the pulse, or between T_(I) and T₂, canbe estimated by sampling Paw, {dot over (V)} and V at multiple pointsprior to, or after, the pulse and estimating ΔPmus/Δt by suitableregression analysis. The standard equations for linear and non-linearregression can be applied to the multiple data sets, to obtain anestimate of Paw, {dot over (V)} and V at T_(I). These are then insertedat the appropriate locations in equations 8 and 8a.

[0073] 6) Use of positive flow pulses (transients): Whereas there isdescribed above the application of the procedure of the invention usingnegative Paw and flow transients (for example, FIG. 1), the sameapproach can be applied to imposed positive flow and Paw transients.Here, Paw, {dot over (V)} and volume are also measured immediatelybefore the perturbation (T_(O)), at or near the point of maximum Paw (orflow) of the positive pulse (T_(I)) (as opposed to the trough of thenegative pulse) and at a third point, either before T_(O), to implementthe extrapolation technique, or after T₁, to implement the interpolationor back extrapolation techniques. The values of Paw, {dot over (V)} andV obtained at the three points (T_(O), T₁ and T⁻¹ or T_(O), T_(I) andT₂) are then inserted in equation 8 or 8a (for extrapolation approach),8 inter or 8a inter (for interpolation approach) or 8 (bextra) and 8a(bextra) (for the back extrapolation approach). Regression analysis canalso be used on multiple data prior to T_(O). In my experience, negativepulses provide more reliable results and are, therefore, preferred. Themore reliable result using negative pulse is likely related to twofactors. First, negative pulses dictate the occurrence of a point atwhich ΔPaw/Δt is zero, which can be used as T_(O) (see FIG. 1). Withpositive pulses, this cannot be assured. There are advantages to makingthe measurements at points where ΔPaw/Δt and Δ{dot over (V)}/Δt are nearzero, (as discussed above). Second, in many patients there aresubstantial differences in time of occurrence of peak flow and peak Pawwhen positive pulses are given, which introduces uncertainty in theresults.

DESCRIPTION OF PREFERRED EMBODIMENT

[0074]FIG. 2 shows an overview of a preferred embodiment of apparatusfor carrying out the present invention. This preferred embodimentreflects the actual system used to validate the inventive procedures ofthe invention in 67 ventilator-dependent patients. The preferredembodiment has several components. Although in FIG. 2, these componentsare shown as distinct from each other, such representation is for thesake of illustration of these components, in actual practice all threecomponents can be incorporated within a single unit (the ventilator).

[0075] A gas delivery unit 10 is a ventilator system that is capable ofdelivering proportional assist ventilation (PAV). A variety ofmechanical systems can be used to deliver PAV and some are commerciallyavailable, which use blower-based, piston-based and proportionalsolenoid systems. PAV is described in U.S. Pat. No. 5,107,830 (Younes),assigned to the assignee hereof and the disclosure of which isincorprorated herein by reference. The ventilator illustrated in thepreferred embodiment consists of a piston 12 reciprocating within achamber 14. The piston 12 is coupled to a motor 16 that applies force tothe piston 12 in proportion to input received from the PAV pressurecontrol unit 18. A potentiometer 20 measures the piston displacementwhich corresponds to the volume change during the ventilator cycle.After certain corrections related to leaks and gas compression, thissignal conveys the amount of gas (volume) received by the patient. Thevolume signal (V) is differentiated using a suitable differentiator 22to result in a flow signal ({dot over (V)}). The PAV pressure controlunit 18 generates a signal that is the sum of a suitably amplified flowsignal and a suitably amplified volume signal with amplification factorsbeing set by the user, which signal is used to control the motor 16. Thepiston chamber 14 receives suitable gas mixture through an inlet port 24and delivers gas to the patient through an outlet port. Duringinspiration, an exhalation valve control circuit closes the exhalationvalve 26 ensuring that the gas pumped by the piston 12 is delivered tothe patient through valve 27. At the end of the inspiratory cycle, theexhalation valve control circuit opens the exhalation valve 26 to allowexpiratory flow to occur prior to the next cycle.

[0076] A micro controller 28 receives the flow and airway pressuresignals. These can be obtained directly from ventilator outputs of flow({dot over (V)}) out and airway pressure (P). Alternatively, flow andairway pressure are measured independently by inserting apneumotachograph (30) and an airway pressure outlet between the Yconnector 32 and the patient. The latter approach is the one illustratedin FIG. 2. Pressure transducers are provided (FT and PT) to generatesignals in proportion to airflow and airway pressure near theendotracheal tube 34. Although this is a more direct way of estimatingpatient flow and airway pressure, reasonably accurate estimates can beobtained from sensors within the ventilator body, remote from thepatient, after allowances are made for tube compression. The patientflow and airway pressure signals are continuously monitored by the microcontroller 28. At random intervals, electric pulses are generated by themicro controller and are conveyed to the PAV delivery system viasuitable output ports. These pulses may be either negative or positive,as described above. The pulse output is connected to the PAV pressurecontrol unit 18 within the PAV delivery system via 36. The electricalpulse results in a temporary decrease or increase in the output of thePAV pressure control unit relative to the output dictated by the PAValgorithm. This, then, results in either a corresponding decrease orincrease in airway pressure for a brief period (for example,approximately 0.2 sec), at a time determined by the micro controller, inselected breaths.

[0077] The basic components of the micro controller 28 used in thispreferred embodiment are shown in FIG. 3. The flow and airway pressuresignals are passed to signal conditioning circuits ((LM324OP-AMPS) orequivalent) which condition the input voltage signals into 0 to +5volts. The two signals are passed through to an analog to digitalconvertor on the micro controller. The digitized flow signal isintegrated to provide inspired volume. A clock circuit allows flow,pressure and volume to be sampled at precise intervals. The basiccomputer is an MC68HC16 with AM29F010 ROM and KM68-1000 RAM. A preferredembodiment of the master computer program includes several functions asfollows:

[0078] (1) A function to identify the beginning of inspiration.Inspiration is deemed to have started when inspiratory flow exceeds acertain threshold (e.g. 0.1 l/sec) and remains above this level for aperiod of at least about 150 msec beyond this point.

[0079] (2) A random number generator function generates a number between4 and 11 which determines the number of breaths between any twosuccessive perturbations. This results in an average of 6 unperturbedbreaths between any two successive perturbations. Any other convenientintegers and average may be chosen. The average number also can beover-ridden by the user through a manual input via a key pad. The usermay elect to deliver the perturbations at a faster average rate to speedup the data collection or, conversely, the frequency of application ofperturbations can be slowed down, as, for example, when the clinicalcondition is fairly stable. Clearly other methods of ensuring thatpulses are applied at random intervals are possible. Pulses may also beapplied at regular intervals, although this may result in anticipatoryresponses by the patient which may corrupt the measurements under somecircumstances.

[0080] (3) An event processor function which controls the time ofapplication and characteristics of the pulse. The timing is adjustedautomatically so that the pulses are delivered in the first half of theinflation phase based on the prevailing duration of the inflation phaseobtaining in previous breaths. The shape of the applied pulse is alsoadjusted automatically to result in a reasonably flat segment in Paw andflow during the pulse near T_(I) (see FIG. 1). The information producedby the event processor is conveyed to the pulse generator (DAC-08, FIG.3) which generates a pulse of about 0.2 second duration or any otherconvenient duration. A pulse invertor and gating circuit (LM660 OP-MPSand CD4052 analogue multiplexor, FIG. 3) is used to produce either apositive pulse or negative pulse.

[0081] (4) A subprogram that causes the values of flow, volume, andairway pressure, sampled at about 6 msec or othre convenient timeinterval, to be stored in data memory over the entire period ofinspiratory flow in breaths receiving pulses.

[0082] (5) A subprogram that scans the above data to determine the timeat which peak Paw occurred prior to the negative deflection (T_(O)), atime about 100 msec or other convenient time interval prior to T_(O)(T_(−I)), the time of occurrence of minimum Paw during the pulse (T_(I))and the time of highest Paw in the post-pulse phase (T₂).

[0083] (6) A subprogram to tabulate values of Paw, {dot over (V)}, andvolume at these four time points for each pulsed breath.

[0084] (7) A subprogram that deletes data points that fall outside thenormal variability of the data. This subprogram also identifies breathssubjected to a pulse perturbation where certain criteria are not met.Data related to these observations are deleted from the tables.

[0085] (8) A program that determines the amplitude of pulses to bedelivered. This is an iterative program. The pulse generator isinitially instructed to deliver negative pulses of small amplitude. Thedecrease in flow during these pulses is noted. If the trough of the flow(i.e. {dot over (V)} @T_(I)) is above about 0.2 l/sec or otherconvenient threshold value, the amplitude of the next negative pulse isincreased and the trough in flow is again noted. Progressive increase inthe amplitude of consecutive negative pulses continues until the troughfalls at approximately 0.2 l/sec or other selected threshold value. Theamplitude of the pulses is then kept constant but the trough flow ismonitored each time. Should the trough rise above 0.2 l/sec or otherselected threshold value and remain elevated for a number of pulses, theamplitude of the pulse is increased again. Conversely if the troughresults in zero flow with resetting of respiratory cycle, the amplitudeof the pulse is decreased. The intent of this subprogram is to maintainthe amplitude of the negative pulses such that the trough in flow duringthe negative pulses is close to, but not zero.

[0086] (9) A subprogram that causes early data to be deleted as new dataare acquired, leaving only the results of a specified number of pulses(e.g. last 20 pulses) in the tables.

[0087] (10) A statistical subprogram to calculate the values ofrespiratory system resistance (R) from equations 8, 8a, 8 inter, 8ainter, 8 (bextra) and 8a (bextra) described above. These derivations maybe obtained from the average values of flow, volume and airway pressuretabulated for negative or positive pulses, as described in detail above.

[0088] (11) A function which results in the display of the results ofdetermined resistance (R) on an LCD or other display, as required.

[0089] Whereas in the embodiment described above, a free-standing microcontroller is illustrated, the same functions performed by this microcontroller can be incorporated into a resident computer within theventilator by suitable programme. It is also recognized that theapplication of this technology is not limited to use with the specificpiston-based PAV delivery system used in the above preferred embodiment.All commercial ventilators suitable for use in the Intensive Care Unitare capable of providing outputs related to flow and airway pressure andthose commercial products which have PAV delivery capabilitiesnecessarily have circuitry or micro controllers that execute the PAValgorithm and which can be interfaced with the automated mechanicsmeasurement system provided herein. Understandably, the system describedabove may have to be changed appropriately to adapt to differentfeatures in various PAV delivery systems, but any such modificationsrequired would be well within the skill of anyone experienced in theart. It is also evident that microprocessors and electronic accessoriesother than those described in the preferred embodiment can be utilizedto accomplish the same objectives.

[0090] It is also recognized that modifications to the algorithmsdescribed above with respect to the preferred embodiment are possible.These include, but are not limited to, the following:

[0091] 1) Using pulse durations that are smaller or longer than 200milliseconds.

[0092] 2) Using positive pressure pulses instead of negative pressurepulses or use of both positive and negative pulses.

[0093] 3) The use of complex pulse forms, for example but not limitedto, bi-phasic pulses instead of unimodel pulses.

[0094] 4) More than one pulse is applied during a given breath.

[0095] 5) Where transient increases or decreases in applied pressure forthe sake of determining resistance are produced by transiently changingthe gain of the PAV assist.

[0096] 6) Where transient perturbations in pressure and flow areproduced by a mechanical system independent of the ventilator itself andincorporated in the external tubing.

[0097] 7) Where transient perturbations in pressure and flow, for thesake of determining resistance, are applied during modes other than PAV,including volume cycled assist, CPAP mode, pressure support ventilationor airway pressure release ventilation, whereby perturbations areproduced by superimposing positive and/or negative transients to theusual control signal of the relevant mode.

[0098] 8) Provision to store the resistance results over extendedperiods of time to be made available for later display to providetime-related trends in such relationships.

[0099] 9) Provision of appropriate circuitry or digital means to effectautomatic changes in the magnitude of flow assist in the PAV mode orassist level in other modes, as the resistance values change (i.e.closed loop control of assist level).

[0100] 10) Where resistance is computed from values obtained from singlepulses as opposed to averages of values obtained from several pulses.

[0101] 11) Where the behavior of Pmus during the pulse is calculated byinterpolation between values at T_(O) and values beyond T_(I) (as perequation 8 inter and 8a inter) as opposed to the preferred method ofextrapolation of data between T_(−I) and T_(O) (as per equation 8 and8a).

[0102] 12) Where the behavior of Pmus during the pulse is calculated bybackextrapolation of values occurring between T_(I) and a point beyondT_(I), as per equations 8 (bextra) and 8a (bextra).

[0103] 13) Where resistance is calculated both by the extrapolationtechnique (equation 8 or 8a) and interpolation technique (equation 8inter and 8a inter) and the result is given as an average, orderivative, of the results of the two methods of calculation.

[0104] 14) Where flow is maintained nearly constant for a period of timebeyond T_(I) instead of allowing it to rise again.

[0105] 15) When the assist is terminated immediately after T_(I).

SUMMARY OF DISCLOSURE

[0106] In summary of this disclosure, the present invention providesmethod and apparatus to determine respiratory resistance (R) duringassisted ventilation of a patent in a unique and simplified manner.Modifications are possible within the scope of the invention.

What I claim is:
 1. A method of determining respiratory systemresistance (R) in a patient receiving gas from a ventilatory assistdevice (ventilator), comprising: estimating the flow rate ({dot over(V)}) and volume (V) of gas received by the patient from the ventilator;estimating pressure near the airway of the patient (Paw); generating asignal that results in a step decrease (negative pulse) in the pressureand/or flow output of the ventilator during selected inflation cycles;measuring Paw, {dot over (V)} and V at a point (T_(O)) near thebeginning of the pulse (Paw_(O), {dot over (V)}, V_(O)), at a point (T₁)near the trough of the negative pulse (Paw₁, {dot over (V)}₁, V₁), andat a point (T⁻¹) preceding T_(O) but after the onset of inspiratoryeffort (Paw⁻¹, {dot over (V)}⁻¹, V⁻¹); and calculating the value ofresistance (R) from the differences between Paw, {dot over (V)} and V atT_(O) and at T₁ and where the change in patient generated pressure(Pmus) in the interval T_(O)→T_(I) (ΔPmus ( T_(O)→T_(I))) is estimated,by extrapolation, from the differences between Paw, {dot over (V)} and Vat T_(O) and at T⁻¹, in accordance with equation (8).
 2. The method ofclaim 1 wherein the estimating by extrapolation step is modified byestimating ΔPmus (T_(O)→T_(I)) from the differences between Paw, {dotover (V)} and V values obtained at two time points preceding T_(O), asopposed to T_(O) and a single preceding time point (T⁻¹).
 3. The methodof claim 1 wherein the estimating by extrapolation step is modified byestimating ΔPmus (T_(O)→T₁) using regression coefficients obtained fromregression analysis of Paw, {dot over (V)} and V values measured atmultiple (>2) points preceding the pulse.
 4. The method of claim 1wherein the estimating step to estimate ΔPmus (T_(O)→T₁) is modified byestimating by interpolation, from the differences between Paw, {dot over(V)} and V values obtained at T_(O) and at a second point (T₂) beyondT_(I), in addition to, or instead of, extrapolation of differencesbetween values at T_(O) and T⁻¹.
 5. The method of claim 1 or 3 whereinΔPmus (T_(O)→T_(I)) is estimated by back extrapolation of valuesobtained beyond T_(I).
 6. The method of any one of claim 1 to 5 whereinthe single R value in equation (8) is replaced by mathematical functionsthat allow for non-linear pressure-flow relations.
 7. The method ofclaim 6, wherein said mathematical function is given by R=K₁+K₂ {dotover (V)}, wherein K1 and K2 are the coefficients defining thenon-linear pressure-flow relation.
 8. The method of claim 7, wherein K₂is replaced by a known or assumed K₂ value of the endotracheal tube ofthe patient.
 9. The method of claim 6 or 7 wherein the coefficientsdefining the non-linear pressure-flow relation (K₁, K₂) are obtained byregression analysis performed on the results of two or more pulsesapplied in separate breaths.
 10. The method of any one of claim 1 to 8,wherein a default elastance value (E) is used in lieu of an actuallymeasured elastance value for the sake of computing differences inelastic recoil pressure between T₀, T_(I) and T_(−I) in equation (8).11. The method of any one of claims 1 to 10, wherein positive pulses aredelivered instead of, or in addition to, negative pulses and the T_(I)values of Paw, {dot over (V)} and V are measured at or near peak Paw orflow of the positive pulse.
 12. The method of any one of claims 1 to 11,including automatically adjusting the amplitude of the pulse dependingon the response to previous pulses.
 13. The method of any one of claims1 to 12, including automatically adjusting the timing of pulseapplication during the inflation phase.
 14. The method of any one ofclaims 1 to 13, including automatically adjusting the shape of the pulseto insure the presence of a flat segment in the Paw/flow signal duringthe pulse for use in measuring T_(I) values.
 15. The method of any oneof claims 1 to 14, wherein the pulses are delivered at random intervals.16. The method of any one of claims 1 to 15 including user selecting oneor more pulse characteristics.
 17. The method of any one of claims 1 to16, wherein the resistance results (R) are reported as averages of theresults of several pulses.
 18. The method of any one of claims 1 to 17,wherein resistance results (R) are stored over time to permit thedisplay of time dependent trends.
 19. The method of any one of claims 1to 18 including identifying pulses with results that fall outside thenormal variability of the data as determined from several data samplingand excluding the results of these pulses from analysis anddetermination of R.
 20. The method of any one of claims 1 to 19including deleting early data as new data are acquired and reporting theresults of the determination of R for a specified number of pulses. 21.The method of claim 20 wherein the specified number of pulses isselected, either by a user or, in the absence of user input, as adefault value (e.g. 20).
 22. The method as claimed in any one of claims1 to 21, wherein the step decrease or increase in Paw or {dot over (V)}is produced by an electromechanical system attached to the externaltubing of the ventilator as opposed to directly interfacing with theventilator control system.
 23. The method of any one of claims 1 to 22,wherein the results of the resistance values are used in closed loopcontrol of an assist level provided by the ventilator.
 24. An apparatuswhich interfaces with ventilatory assist devices (ventilators)determining respiratory system resistance (R), comprising: a flowmeter,with associated electronic circuitry, that estimates the flow rate ({dotover (V)}) and volume (V) of gas received by a patient; a pressuresensor that estimates pressure near the airway of the patient (Paw); andelectronic circuitry which receives the Paw, {dot over (V)} and Vsignals from above mentioned circuitry and which is also connected tothe control system of the ventilator, comprising: circuitry thatgenerates an output that results in a step decrease (negative pulse) inthe pressure and/or flow output of the ventilator during selectedinflation cycles; circuitry that measures Paw, {dot over (V)} and V at apoint (T_(O)) near the beginning of the pulse (Paw_(o), {dot over (V)}o,V_(O)), at a point (T_(I)) near the trough of the negative pulse (Paw₁,{dot over (V)}₁, V₁), and at a point (T₁) preceding T_(O) but after theonset of inspiratory effort (Paw⁻¹, {dot over (V)}⁻¹, V⁻¹); circuitry tocalculate the value of resistance (R) from the differences between Paw,{dot over (V)} and V at T_(O) and at T_(I) and where the change inpatient generated pressure (Pmus) in the interval T_(O)→T_(I) (ΔPmus(T_(O)→T_(O))) is estimated, by extrapolation, from the differencesbetween Paw, {dot over (V)} and V at T_(O) and at T⁻¹, in accordancewith equation (8).
 25. The apparatus of claim 24 wherein ΔPmus(T_(O)→T_(I)) is estimated, by extrapolation, from the differencesbetween Paw, {dot over (V)} and V values obtained at two time pointspreceding T_(O), as opposed to T_(O) and a single preceding time point.26. The apparatus of claim 24 wherein ΔPmus (T_(O)→T_(I)) is estimated,by extrapolation, using regression coefficients obtained from regressionanalysis of Paw, {dot over (V)} and V values measured at multiple (>2)points preceding the pulse.
 27. The apparatus of claim 24 wherein ΔPmus(T_(O)→T_(I)) is estimated, by interpolation, from the differencesbetween Paw, {dot over (V)} and V values obtained at T_(O) and at asecond point (T₂) beyond T_(I) in addition to, or instead of, as opposedto extrapolation of differences between values at T_(O) and T⁻¹.
 28. Theapparatus of claim 24 or 26 wherein ΔPmus (T_(O)→T_(I)) is estimated byback entrapolation of values obtained beyond T_(I).
 29. The apparatus ofany one of claims 24 to 28 wherein the single R value in the equationsis replaced by mathematical functions that allow for non-linearpressure-flow relations.
 30. The apparatus of claim 29 wherein saidmathematical function is given by R=K₁+, wherein K1 and K2 arecoefficients defining the non-linear pressure-flow relation.
 31. Theapparatus of claim 30 wherein K₂ is replaced by a known or assumed K₂value of the endotracheal tube of the patient.
 32. The apparatus ofclaim 29 or 30 wherein the coefficients defining the non-linearpressure-flow relation (K₁, K₂) are obtained by regression analysisperformed on the results of two or more pulses applied in separatebreaths.
 33. The apparatus of any one of claims 24 to 31 wherein adefault elastance value (E) is used in lieu of an actually measuredelastance value for the sake of computing differences in elastic recoilpressure between T_(O), T₁ and T_(−I).
 34. The apparatus of any one ofclaims 24 to 33 wherein positive pulses are delivered instead of, or inaddition to, negative pulses and the T_(I) values of Paw, {dot over (V)}and V are measured at or near peak Paw or flow of the positive pulse.35. The apparatus of any one of claims 24 to 34 including algorithms toautomatically adjust the amplitude of the pulse depending on response toprevious pulses.
 36. The apparatus of any one of claims 24 to 35including algorithms which automatically adjust the timing of pulseapplication during the inflation phase.
 37. The apparatus of any one ofclaims 24 to 36 including algorithms that automatically adjust the shapeof the pulse to insure the presence of a flat segment in the Paw/flowsignal during the pulse for use in measuring T_(I) values.
 38. Theapparatus of any one of claims 24 to 37 wherein the pulses are deliveredat random intervals.
 39. The apparatus of any one of claims 24 to 38including a user interface that permits the user to select one or morepulse characteristics.
 40. The apparatus of any one of claims 24 to 39wherein the resistance results are reported as averages of the resultsof several pulses.
 41. The apparatus of any one of claims 24 to 40wherein resistance results are stored over minutes, hours or days topermit the display of time dependent trends.
 42. The apparatus of anyone of claims 24 to 41 including algorithms which identify pulses withresults that fall outside the normal variability of the data and whichexclude the results of these pulses from analysis.
 43. The apparatus ofany one of claims 24 to 42 including algorithms which delete early dataas new data are acquired reporting the results of a specified number ofpulses.
 44. The apparatus of claim 43 wherein the specified number ofpulses is selected by user or, in absence of user input, as a defaultvalue (e.g. 20).
 45. The apparatus of any one of claims 24 to 44 whereinsome or all necessary components are incorporated within the main bodyof the ventilator.
 46. The apparatus of any one of claims 24 to 45wherein the step decrease or increase in Paw or {dot over (V)} areproduced by an electromechanical system attached to the external tubingof the ventilator as opposed to directly interfacing with the ventilatorcontrol system.
 47. The apparatus of any one of claims 24 to 46 whereinthe results of the resistance values are used in closed loop control ofthe assist level provided by the ventilator.